Image-forming MR methods which utilize the interaction between magnetic fields and nuclear spins in order to form two-dimensional or three-dimensional images are widely used nowadays, notably in the field of medical diagnostics, because for the imaging of soft tissue they are superior to other imaging methods in many respects, do not require ionizing radiation and are usually not invasive.
According to the MR method in general, the body of the patient to be examined is arranged in a strong, uniform magnetic field (B0 field) whose direction at the same time defines an axis (normally the z-axis) of the co-ordinate system on which the measurement is based. The magnetic field produces different energy levels for the individual nuclear spins in dependence on the magnetic field strength. Transitions between these energy levels can be excited (spin resonance) by application of an electromagnetic alternating field (RF field, also referred to as B1 field) of defined frequency (so-called Larmor frequency, or MR frequency). From a macroscopic point of view the distribution of the individual nuclear spins produces an overall magnetization which can be deflected out of the state of equilibrium by application of an electromagnetic pulse of appropriate frequency (RF pulse), so that the magnetization performs a precessional motion about the z-axis. The precessional motion describes a surface of a cone whose angle of aperture is referred to as flip angle. The magnitude of the flip angle is dependent on the strength and the duration of the applied electromagnetic pulse.
After termination of the RF pulse, the magnetization relaxes back to the original state of equilibrium, in which the magnetization in the z direction is built up again with a first time constant T1 (spin lattice or longitudinal relaxation time), and the magnetization in the direction perpendicular to the z direction relaxes with a second time constant T2 (spin-spin or transverse relaxation time). The variation of the magnetization can be detected by means of one or more receiving RF coils which are arranged and oriented within an examination volume of the MR device in such a manner that the variation of the magnetization is measured in the direction perpendicular to the z-axis.
In order to realize spatial resolution in the body, linear magnetic field gradients extending along the three main axes are superposed on the uniform magnetic field, leading to a linear spatial dependency of the spin resonance frequency. The signal picked up in the receiving coils then contains components of different frequencies which can be associated with different locations in the body. The MR signal data obtained via the RF coils corresponds to the spatial frequency domain and is called k-space data. A set of k-space data is converted to a MR image by means of Fourier transformation or other appropriate reconstruction algorithms.
A significant level of acoustic noise occurs in conventional MR imaging sessions which is caused by mechanical oscillations of the magnetic field gradient coils placed in the static main magnetic field. The Lorentzian forces induced when an electrical current is applied to the gradient coils make them physically move. This displacement is dependent on the strength of the static magnetic field, the amplitude of the electrical current applied, and the frequency and waveform of the magnetic field gradient switching. The amplitude of the acoustic noise within the examination volume of a MR imaging device typically varies from 94 to 135 dB depending on various parameters: the hardware characteristics of the respective MR imaging device, the extent of vibration of the system, the type and the parameters (for example the repetition time) of the MR imaging sequence applied, the number of slices acquired, etc. The high level of acoustic noise induces stress and discomfort in the examined patients. Hearing protection is required to prevent hearing impairment.
The drawbacks of acoustic noise can be overcome by recently developed virtually silent MR imaging techniques, in which RF excitation as well as acquisition of MR signals are performed in the presence of a magnetic field gradient. The magnetic field gradient is applied for purely frequency-encoded, radial centre-out k-space encoding. In these known approaches, the spatially non-selective excitation must uniformly cover the full frequency bandwidth spanned by the readout magnetic field gradient, which is typically accomplished by radiating short, hard RF pulses. The acquisition of a free induction decay (FID) signal starts immediately after radiation of the RF pulse. After the FID readout, only minimal time is required for setting of the next readout magnetic field gradient before the next RF pulse can be applied, thus enabling very short repetition times (TR). The magnetic field gradient vector determining the readout direction is incrementally varied from repetition to repetition until a spherical volume in k-space is sampled to the required extent. Such radial centre-out k-space scanning techniques are sometimes referred to as “koosh ball”-scanning, with the radial k-space “spokes” and their arrangement in k-space resembling the filaments (strings) of the known toy ball design. Without the need for switching off the readout magnetic field gradient during the whole scan, MR imaging can be performed virtually silently (see, for example, Weiger et al, Magnetic Resonance in Medicine, vol. 70, p. 328-332, 2013). Further, the U.S. Pat. No. 5,570,018 mentions spatial encoding of magnetic resonance signal by way of sinusoidally varying magnetic gradient fields in two orthogonal directions (y and z) and using periods that differ by a factor of two.
Intrinsically, the known silent radial centre-out k-space encoding techniques deliver only a proton-density weighted contrast of the reconstructed MR image.